1. Field of the Invention
The present invention is directed to a method for detecting an ultrasound contrast agent in a soft tissue and quantitating blood perfusion through regions of tissue by detecting the contrast agent in the tissue.
2. Description of the Related Art
An ultrasound contrast agent is a solution of small gas bubbles (diameter xcx9c5 xcexcm) that is injected into the blood stream. Such bubbles show strong and non-linear scattering of ultrasound at the frequencies used for medical ultrasound imaging. Medical applications of the contrast agents include, but are not limited to, enhancing imaging of blood vessels, improving the detection of the endocardium as a border of the ventricular cavities, and improving the detection of blood jets through leaking cardiac valves or septal defects.
There has also been great hopes that ultrasound contrast agent should be able to detect and quantify blood perfusion through the tissue, especially the myocardial tissue where coronary disease strongly influences the myocardial perfusion. The widespread occurrence of coronary artery disease as a major cause of death in the western world has made this application of the contrast agents and methods for detection of the contrast agents in the tissue a target for development of various types of ultrasound contrast agents.
Second harmonic ultrasound imaging, is today the commonly used method for detecting and imaging ultrasound contrast agent in the tissue. The non-linear elastic properties of the contrast agent bubbles produce higher harmonic components and sub harmonic components of the transmitted pulse frequency band in the scattered signal directly in the scattering process. However, with the present wideband transducer technology it is only possible to utilize the second harmonic component of the signal by transmitting an ultrasound pulse with frequency spectrum in the lower part of the active frequency band of a wideband transducer. The second harmonic component of the scattered signal is then received in the upper part of the transducer frequency band.
Forward propagation of the ultrasound pulse through the tissue produces a distortion of the pulse due to pressure dependent propagation velocity. The distortion is limited by acoustic power absorption in the tissue, so that in practice we get high enough amplitude of the 2nd harmonic band in the pulse to use this harmonic band for imaging of soft tissue itself. With venous injection of contrast agent, the 2nd harmonic frequency band of the scattered signal from the myocardial tissue has, however, comparable amplitude to the 2nd harmonic frequency band of the signal from contrast agent in the myocardium. The back scattered tissue signal then represents a background noise for the detection of the contrast agent, and hence limits the detectability of low concentrations of the agent based on the 2nd harmonic component of the back scattered signal.
The non-linear elasticity of the contrast agent is much stronger than that of the tissue. Accordingly, considerable distortion of the scattered pulse directly in the scattering process results with high amplitudes in the 3rd and 4th harmonic component of the incident pulse frequency band. More importantly, the scattered amplitudes from the contrast agent in these frequency bands are much stronger than the scattered amplitudes in the 3rd and 4th harmonic frequency bands from the tissue. Therefore, the use of harmonic frequency bands higher than the 2nd component of the transmitted pulse frequency band for detection and imaging of the contrast agent provides improved separation between the signal amplitudes from the contrast agent and the signal amplitudes from the tissue.
There are however several practical problems in utilizing the 3rd and 4th harmonic component of the transmitted frequency band for detection and imaging of ultrasound contrast agent, as well as using such imaging to grade the degree of blood perfusion through tissue:
The first problem is that the present medical ultrasound transducers have so narrow a bandwidth that it is not possible to transmit a pulse with frequency band around f0, and receive back-scattered frequency components in the frequency bands around 3f0 and 4f0 with adequate sensitivity. Adequate wideband transducers have been made by highly damping the transducers, but this reduces the sensitivity to the signal scattered from the contrast agent in the myocardium below tolerable limits. According to the present invention, a transducer is used with capabilities of transmitting frequencies in a band around f0, with high sensitivity in the receive band around the 3rd or 4th harmonic component of the transmit band. In the particular implementation of the invention, the high receive sensitivity is obtained by using resonant operation of the transducer in the receive band with minimal dampening.
For the invention to fully work, the transmitted pulse must have sufficiently limited amplitude in the receive frequency bands. A solution to this problem is presented according to the invention by either bandpass filtering the transmitted pulse both in the transducer and/or electrically before driving the transducer, or by using band limited pulse generator with linear drive amplifiers of the array transducer elements.
A second problem associated with utilizing the 3rd and 4th harmonic component of the transmitted frequency band for detection and imaging of ultrasound contrast agent is that the pulse distortion in the scattering from the contrast agent bubbles, highly depends on the amplitude of the pulse incident on the bubble. Absorption of the transmitted pulse attenuates the incident amplitude with depth, depending on the transmitted frequency f0. In addition, the beam divergence past the transmit focus will produce an amplitude attenuation with depth. In the normal imaging situation, the amplitude of the transmitted pulse hence attenuates with depth, giving a subsequent reduction in the distortion of the scattered pulse from the contrast agent with depth. This produces a depth variable detection of the contrast agent, and presents severe problems for imaging and quantitation of regions of reduced blood perfusion in the myocardium.
The power absorption in the tissue considerably reduces with the frequency, being approximately 0.5 dB/cmMHz. Hence, by using a low transmitted center frequency at for example f0=0.875 MHz, the total absorption attenuation from 2-10 cm is xcx9c3.5 dB. Geometric focussing of the beam to the far end of the image range may be used to compensate for this absorption attenuation. Due to diffraction at such low frequencies, the maximal amplitude in the focussed beam is found closer to the transducer than the geometric focus. By locating the geometric transmit focus beyond the image range, at for example 12 cm, the transmit beam focussing will give a gain of xcx9c3.4 dB from 2-10 cm with a 18 mm circular aperture. Hence, by using sufficiently low transmit frequency, one can obtain approximately constant incident amplitude over a limited image range by proper selection of transmit aperture and focus.
A transmit center frequency of 0.875 MHz, gives 3rd and 4th harmonic center frequencies at 2.625 MHz and 3.5 MHz, which are typical frequencies used for cardiac imaging. These frequencies produce tolerable absorption attenuation of the backscattered signal so that it can be compensated for by a depth variable receiver gain. A receive frequency in the range of 2.5-4 MHz also gives a lateral resolution of the receive beam comparable to that with regular echocardiography.
To obtain a narrower transmit beam at all ranges and to further improve the equalization of the incident pulse amplitude with depth according to the invention, the depth along the receive beam is divided into sub-ranges. Each sub-range is observed at different time intervals with different transmit pulses, where the focus of each transmit pulse is located within the corresponding receive range, and both the transmit focus, the transmit amplitude, and the transmit aperture are adjusted for optimal equalization of the incident pulse amplitude within the corresponding receive range, under the actual absorption of the ultrasound in the tissue.
To cover a range from, for example, 2 cm to 15 cm with minimal variation of the incident pulse according to the invention, the range could be divided into subranges from 2-6 cm to be interrogated with a transmit pulse with focus at 7 cm with a reduced transmit aperture to secure sufficient focal depth, followed by a 2nd transmit pulse focussed at say 18 cm, using larger transmit amplitude and transmit aperture to achieve the same incident pulse amplitude at 2 and at 15 cm. Further detailed optimization of number of transmit pulses, with corresponding transmit foci, amplitudes, and apertures can be done within the scope of the invention. Pulse destruction by the transmit pulses must also be taken into the account in this optimization, as described below.
A third problem associated with utilizing the 3rd and 4th harmonic component of the transmitted frequency band for detection and imaging of ultrasound contrast agent is that the power of the received signal from a depth, is proportional to the concentration of contrast agent in the tissue. The source of contrast agent in a region of the myocardium is the inflow of blood to the region with a sink of contrast agent produced by the venous outflow. In the stationary situation with some blood flow through the tissue, the concentration of contrast agent in a tissue region is hence given by the product of the blood concentration in the region and the concentration of contrast agent in the inflowing blood. The contrast agent concentration will in this situation not be reduced before close to complete blocking of the blood flow to the region occurs. Therefore, in a stationary situation, the received signal power from a range will give limited quantitative grading of the perfusion through the tissue region.
One method to improve the grading of the perfusion is to introduce an additional sink of contrast agent in the region such, for example, as by destroying the contrast agent with high amplitude incident ultrasound pressure pulses. Partial destruction of the contrast agent will give a concentration that depends on both the inflow rate of blood to the region (i.e., the source of contrast agent) and the destruction rate (i.e., the additional sink of contrast agent). The concentration of contrast agent will in this case quantitatively be reduced with reduced blood perfusion in the tissue. By comparing the signal level from one region with the signal level from regions with normal perfusion, the received signal level will give a quantitative measure of regionally reduced perfusion. With complete destruction of the contrast agent in the image region, one can use the re-filling time of contrast agent in the tissue as a quantitative measure of the perfusion through the tissue.
It is then important that the whole image range is imaged with incident pulses of equal amplitude so that variations in backscattered amplitude are produced by variations in the concentration of contrast agent and not by variations in the incident pulse amplitude. As some destruction occurs at practical image amplitudes, it is also important to design the transmit pulses so that this destruction is the same throughout the whole image range.
Also, when transmitting dedicated destruction pulses for partial destruction of the contrast agent to quantitatively grade the perfusion, one must assure that the degree of destruction is practically constant throughout the whole image range. The amplitude of the destruction pulses can be depth tailored as for the image pulses, using multiple transmit pulses with optimized transmit foci, amplitudes, and apertures. The pulse with focus in one sub-range, will then provide some contrast agent destruction at other ranges, and the whole set of transmit pulses for each image direction, must be designed for equal and limited destruction of the contrast agent along the whole image depth.
Multiple foci destruction pulses will also provide minimal width of the destruction beam for all depths, ensuring that the destruction pulses for one image beam produce limited contrast agent destruction in neighboring image beams. Destruction pulses with narrow focus at small depths are then used first to destroy contrast agent at low ranges. The rapid geometric widening of the beam past the focus will then produce an attenuation of the destruction pulse which reduces contrast agent destruction at larger depths. Using a higher frequency of the shallow range destruction pulses, will also increase the attenuation of these pulses at deeper ranges. The destruction of the contrast agent at deeper ranges, is then followed up with new destruction pulses with deeper foci, larger transmit amplitudes and apertures, and possible lower center frequency. The overlap of destruction from the pulses at all ranges must be taken into account. Therefore, the whole set of destruction pulses must be designed for each receive beam direction so that even destruction of the contrast agent over the whole image range occurs.
Contrast agent destruction by neighboring beams will only complicate the imaging at the edges of the scan, where the edge beam has a single neighbor. Further into the image, all beams will get the same destruction by the neighboring beams.
A method for detection of ultrasound contrast agent in soft tissue according to the present invention includes utilizing an ultrasound transmit beam former and transducer array assembly for transmitting directive, focussed ultrasound pressure pulses with steerable transmit amplitude, transmit aperture, transmit focus, and transmit direction, and with temporal frequency components within a limited band B centered at frequency f0, towards a region of soft tissue that contains ultrasound contrast agent bubbles. The transmit pulse parameters are arranged, possibly using multiple transmit pulses, so that the incident pressure pulse that is utilized for imaging of the contrast agent for a particular depth, has minimal variation over the actual image range. The non-linearly distorted, back-scattered ultrasound signal is received from both the tissue and the ultrasound contrast agent bubbles with the same ultrasound transducer assembly and the received array element signals are passed through a receiver beamformer that has a steerable spatially directive receiver sensitivity.
The transducer assembly has high sensitivity at the receive band of frequencies centered at 3f0 and/or 4f0 for maximal sensitivity of the distorted, non-linearly scattered signal from the contrast agent bubbles. The received signal is high-pass filtered so that the lowest frequency component of the resulting signal is at least 2 times higher than the frequency component of the transmitted signal. The amplitude of the high-pass filtered signal is used for detecting ultrasound contrast agent bubbles buried within the tissue, and for imaging of contrast agent bubbles in the tissue.
The depth variation of the amplitude of the incident pressure pulse is minimized by positioning the transmit focus deeper than the image range.
The width of the incident beam at each location is reduced, and the depth variation of and amplitude of the incident pressure pulse is minimized by dividing the total imaged depth range into sub-ranges, where a separate transmit pulse is used to interrogate each sub-range consecutively in time, arranging the transmit focus, the transmit aperture, and the transmit amplitude for each pulse so that the pressure pulse amplitude incident on the contrast agent bubbles at their location in the absorbing tissue is practically equal for each sub range.
The transmitted center frequency f0 can be less than 1 MHz in the preferred embodiment.
To improve the sensitivity of the receiving transducer assembly in the receive band, a backing mount of the transducer with characteristic acoustic impedance less 30% of that of the active electro-acoustic layer may be used. Alternatively, the improved sensitivity of the receiving transducer assembly in the receive band may be facilitated by using a backing mount of the transducer with characteristic acoustic impedance greater than 150% of that of the active electro-acoustic layer.
The sensitivity of the receiving transducer assembly in the receive band is also facilitated by making the transducer assembly resonant in this band.
The ultrasound transducer array comprises an electro-acoustic active layer divided into several transducer elements with a front and a back face, a 1st thin electrode layer covering the front face, and a 2nd thin electrode layer covering the back face. The electrodes are electrically connected to electric terminals for coupling of energy between the electric terminals and acoustic vibrations in the transducer elements.
A substrate layer is mounted on the back side of the acoustic layer with approximately the same acoustic properties as the active layer. The back layer is mounted on an acoustically absorbing backing with acoustic impedance much lower than the two layers.
The ultrasound transducer array further comprises at least one acoustic matching layer mounted on the front face of the active layer and acoustically in contact with the tissue. The acoustic properties and thicknesses of the matching layers are adjusted to facilitate improved acoustic power transfer to and from the tissue and to facilitate a wide bandwidth of the electro-acoustic transfer function to transmit a band-limited ultrasound pulse centered at f0 into the tissue, and to receive backscattered ultrasound pulses in the 3rd or 4th harmonic component, or both, of the transmit band. The substrate layer also is electro-acoustically active and divided into individual transducer elements with common faces to the first transducer elements, with a third, thin electrode layer on the back face of the elements, which can be combined with the 2nd or the 1st electrodes for coupling of energy between the electric terminals of the electrodes and acoustic vibrations in the transducer combined transducer elements.
Two of the 3 electrode layers may be connected to the transmit amplifiers to transmit the low frequency acoustic pulse, and another two of the 3 electrode layers may be coupled to the receiver amplifiers to receive the back scattered acoustic energy from the contrast agent bubbles.
The method for detecting the contrast agent may be used for quantitating variations in tissue blood perfusion. To accomplish this, the ultrasound contrast agent in the tissue may first be destroyed uniformly with depth and direction in the tissue with a controllable degree, followed by imaging of the backscattered signal power from contrast agent in the tissue.
Partial destruction of the contrast agent may be performed so that the amplitude of the backscattered signal in the 3rd or 4th harmonic component of the transmit frequency band gives a regional grading of the perfusion.
Separate destruction pulses may be used to controllably destroy the contrast agent uniformly over the whole image field.
The contrast agent may also be fully destroyed in the tissue, and imaging may be performed at a time interval after this destruction, so that the amplitude of the back-scattered signal in the 3rd or 4th harmonic component of the transmit frequency band gives a regional grading of the refilling time of blood into the tissue, and hence the blood perfusion through the tissue.
The timing of the contrast agent destruction may be derived from the electrocardiogram (ECG), and imaging may be performed at a selected period in the cardiac cycle derived from the ECG.
The increase in image intensity is followed for many heart cycles after the contrast destruction to obtain complete re-filling curves of contrast agent into different regions of the tissue, for regional grading of the perfusion into the tissue.
Other objects and features of the present invention will become apparent from the following detailed description considered in conjunction with the accompanying drawings. It is to be understood, however, that the drawings are designed solely for purposes of illustration and not as a definition of the limits of the invention, for which reference should be made to the appended claims. It should be further understood that the drawings are not necessarily drawn to scale and that, unless otherwise indicated, they are merely intended to conceptually illustrate the structures and procedures described herein.